In joint surgery it is common practice today to anchor components of replacement joints by using as bone cement a two-component resin which polymerizes during the operation at normal temperatures and which, on account of its plastic properties leads to an interlocking of the prosthesis component in the bony sheath. Because of its physical properties, the bone cement shrinks onto the prosthesis resulting in a closed metal-to-cement contact.
The bone cements commonly used are polymethylmethacrylate (PMMA) consisting of powdery bead polymers which are superficially dissolved by liquid monomers and embedded during the polymerization process. During mixing the polymer is immersed in the monomers. The PMMA beads are superficially dissolved and embedded in a composite manner. Despite their widespread use PMMA and related bone cements tend to represent the "weak link" in prosthesis fixation.
The long term success of a total joint prosthesis depends on the continued function and interaction of each of the components of the prosthetic system. In a cemented total hip prosthesis, for instance, stress transfer from the pelvis to the femur is a function of the materials between the two bones (e.g. bone-PMMA-metal-Ultra-high Molecular Weight Polyethylene-metal-PMMA-bone) and the interfaces between the materials. The weakest of the materials is the PMMA, with the lowest fracture toughness and ultimate strength.
The common mode of failure of total joint prostheses is aseptic loosening. X-ray examinations of patients with loosened prostheses often reveal a radiolucent line in the bulk of the cement, indicating that the cement has fractured. Because the geometry of the prosthesis is complex, the state of stress is also highly complex, and the reasons for cement failure are not clear. For example, it has been postulated that the integrity of the metal stem/PMMA interface is the critical link in the performance of the prosthesis; however, the cause and effect relationship between the metal prosthesis/PMMA interface failure and cement fracture is not well understood although the fracture mechanics of the two phenomena are most likely linked. The improvement of the fracture characteristics of the bone cement, however, is a problem that has received some attention in recent years.
The composition of the PMMA used for total joint surgeries today is substantially the same as that used 20 years ago; very little has been done to improve the material itself. The acceptable success rate of cemented prostheses was achieved using existing cements, however, in a predominantly elderly patient population and with improved surgical handling techniques. The 90% success rate at ten years is good, but should be improved. Cement failures do occur, and generally lead to revision surgery. Furthermore, younger patients now receiving total joint replacements have a greater life expectancy than the design expectations of the total joint prosthesis. Improvement of the bone cement, exclusively, may not solve every problem associated with total joint replacements. But, by making improvements in each component of a total joint prosthesis, including the cement, the success rate of prostheses will improve, and mechanical failures can be virtually eliminated.
Increasing the longevity of PMMA by improving the resistance to failure of the polymer has received some, albeit surprisingly little, attention in the bioengineering literature in the past ten years. One suggested method of improvement was to formulate a new bone cement, based on n-butyl methacrylate, rather than the methyl methacrylate monomer. It has been reported that the material showed a higher ductility, a higher apparent fracture toughness, and a greater fatigue life. However, the actual fracture toughness determined by separate impact tests showed no improvement of the new cement with respect to PMMA cements. An even more detrimental result was that the new polymer had only half the modulus and half the ultimate tensile strength of PMMA.
Another method of attempting to improve PMMA was the addition of a reinforcing phase, generally short fibers or whiskers. Early work was done by Knoell, et al., Ann. Biomed, Eng., 3, 1975, pp. 225-229 with carbon fibers approximately 6 mm in length, 1, 2, 3 and 10% fiber content by weight (approximately 0.67, 1.33, 1.96 and 5.87% fiber content by volume, with measured increases of 100% in the average Young's modulus for the reinforced PMMA. They also reported a decrease in peak curing temperature of the reinforced PMMA. They found the reinforced cement viscous and difficult to mix, and they altered the ratio of powder polymer to liquid monomer to facilitate mixing of the reinforced cement. Pilliar, et al., J. Biomed. Mater Res., Vol. 10, 1976, pp. 893-906); Fatigue of Filamentary Component Materials ASTM STP 636, eds. Reifsnider, et al., ASTM 1977, pp. 206-227; used carbon fibers (6 mm length, 7 micrometers diameter) with a 2% volume content. They measured a 50% improvement in tension-tension fatigue limit, improved impact performance, and a 36% increase in ultimate tensile strength. However, it was implied that the reinforced PMMA had poor intrusion characteristics due to increased viscosity, and poor fiber distribution. Wright, et al. J. Mater. Sci. Let., 14 1979, pp. 503-505, did preliminary studies using PMMA reinforced with chopped aramid fibers. PMMA reinforced with 5.17% by volume (7% by weight) exhibited a 74% increase in fracture toughness over the plain PMMA. They were not able to produce reinforced PMMA with a fiber content greater than 5% by volume because of mixing and handling difficulties. Beaumont, J. Mater. Sci., 12, 1977, pp. 1845-1852 included glass beads in the PMMA mass and measured a 10.sup.3 decrease in crack propagation velocity, using 30% volume content of the beads.
Very few investigations involved the use of metal fibers to reinforce PMMA. Taitsman and Saha, J. Bone Joint Surg., Vol. 59-A, No. 3, April 1977, pp. 419-425, used large diameter (0.5 to 1.0 mm) stainless steel and vitallium wires as a reinforcing phase. They embedded 1, 2, or 3 wires in their PMMA specimens. They reported up to an 80% increase in tensile strength of the PMMA, with three embedded vitallium wires, but noted that clinical applications of their wire reinforced cement were limited. Taitsman and Saha's use of reinforcing wires is analogous to reinforcing bars embedded in structural concrete, and not a homogeneous, fiber composite material. Fishbane and Pond, Clin. Orthop., No. 128, 1977, pp. 194-199, reinforced industrial grade PMMA and PMMA bone cement with stainless steel whiskers (0.5-1.0 mm length and 65 micron diameter; 3-6 mm length and 90 microns diameter). They determined that the addition of fibers up to 6.5% by volume improved the compressive strength by nearly 100% for the industrial PMMA, but only 25% for the surgical grade PMMA. The compressive strength of PMMA is not a critical property for the longevity of the cement in vivo. These authors postulate that the reason for the decreased performance of the surgical PMMA was: ". . . due to the limitations imposed by the (surgical) methacrylate preparation technique."
Schnur and Lee, J. Biomed. Mater Res., Vol. 17, 1983, pp. 973-991, used titanium (Ti) sheet, wire, mesh and powder as a reinforcing phase with the purpose of increasing the modulus of PMMA to the modulus of cortical bone. A 16% volume fraction of 1 mm diameter wires (a total of 25 wires) increased the modulus of the PMMA by 380%, and the maximum compressive stress by 75%. The concept is again similar to the reinforcing bars embedded in concrete.
The more recent work in reinforcing PMMA bone cement as reported in the literature, has involved either carbon, graphite, or aramid fibers. Robinson, et al., J. Biomed. Mater Res., Vol. 15, 1981, pp. 203-205, tested both regular PMMA and low viscosity PMMA cement (available from Zimmer Co., Warsaw, Ind.) reinforced with 2% volume of carbon fibers (1.5 mm in length, 10 microns diameter). Both reinforced cements exhibited an increase in apparent fracture toughness (notched bending strength tests) of approximately 32% over their plain counterparts. Surface fractography revealed no evidence of fiber fracture, indicating that the increases in "toughness" was due principally to fiber pull out. In other work with carbon fiber reinforced PMMA an order of magnitude decrease in crack propagation velocity was attributed to the carbon fiber reinforcement of both the regular and low viscosity cements.
Saha and Pal, J. Biomechanics, Vol. 17, No. 7, 1984, pp. 467-478, tested PMMA reinforced with carbon fibers, 0.67% by volume (1% by weight; 6 mm length, 8 microns diameter) and PMMA reinforced with aramid fibers (Dupont Kevlar-29), 1.61 and 3.82% by volume (2 and 4% by weight; 12-13 mm length, unspecified diameter). The reinforced PMMA showed an increase in the ultimate compressive strength of 20.5% for the carbon fibers, and 19.5% and 28.7% for the 1.61 and 3.82% volume % aramid fibers, respectively. Two important consequences of the addition of fibers to PMMA were proposed: The peak temperature of the reinforced PMMA was lower than the plain PMMA, and the addition of fibers changed the workability of the cement. They recognized that uniform dispersion of fibers was not achieved. Saha and Pal studied a machine mixing technique for distributing the fibers. Their claim that machine mixed specimens were stronger than non-machine mixed specimens is misleading. They used a different shaped fiber for their machine mixed specimens. It is the superior shape of the fiber which is presumed to account for the increase in strength. Machine mixing was never shown to improve the properties of reinforced PMMA.
Ekstrand, J. Biomed. Mater Res., Vol. 21, 1987, pp. 1065-1080, fabricated carbon fiber reinforced PMMA by using clinically irrelevant, industrial fabrication techniques with fiber content as high as 16.38% by volume (40% by weight).
Recent work by Pourdeyhimi, et al., Ann. Biomed. Eng., 14, 1986, pp. 277-294, studied the effect of the fiber content of the fracture toughness of hand-mixed, reinforced, dental PMMA. They used aramid fibers from 0.82 to 5.17% by volume (1 to 7% by weight), and graphite from 0.67 to 5.87% by volume (1 to 10% by weight). For each type of fiber reinforced cement, the fracture toughness increased with increased fiber content. The aramid fiber specimens showed a greater increase than the carbon fiber specimens of the same weight percent, presumably because the energy dissipated in the micromechanisms of failure is greater for the aramid fibers than for the carbon fibers. They were not able to produce a uniform distribution of the fibers.
U.S. Pat. No. 4,064,566 to Fletcher, et al. discloses a graphite fiber reinforced bone cement of the acrylic type stated to have mechanical properties more nearly matched to those of bone and thermal curing characteristics resulting in a lower exothermic temperature reaction during curing. The bone cement composition is a dispersion of from 2 to 12% by weight of very fine high modulus graphite fibers having a diameter below 50 microns and between 0.1 and 15 mm in average length in a solution of biocompatible polymer dissolved in a reactive monomer. Fletcher reports only an increase in the modulus of the bone cement, which may not be of primary concern to a reinforced bone cement, and indeed can be detrimental to the prosthesis system. There was a decrease in compressive strength, and more negatively, a decrease in flexural strength for the reported composite.
U.S. Pat. No. 4,239,113 to Gross, et al. discloses an acrylic based bone cement filled with between 15 and 75% by weight of inorganic material comprised of about 90 to 99% by weight of a bio-active glass ceramic powder and about 1 to 10% by weight of vitreous mineral, e.g., glass, fibers having a length below about 20 mm. The particle size of the powder is from 10 to 200 micrometers. Fiber diameters are not disclosed. Improvements in impact strength, and compression strength were reported. However, a significant decrease in the bending strength and an increase in the modulus of elasticity were also reported. Further, there are no examples given as to the clinical usefulness of this cement. Bioactive glass degrades with time, and hence the integrity of the reinforced bone cement will also degrade with time. The controlled experimentation shows that there is no mechanical improvement due to the fiber reinforcing phase alone. Any improvement is due to the combination of bioactive glass and fiber in concert. Since the bioactive glass degrades with time, the properties of the reinforced cement proposed by Gross, et al. will also degrade with time. Davidson, in U.S. Pat. No. 4,735,625, reports the invention of a reinforced bone cement formed using a sock-like mesh of a fiber-like material to reinforce the cement in the vicinity of the prosthesis. The volume of "reinforced" bone cement is limited; critical areas are not reinforced. Draenert, in U.S. Pat. No. 4,365,357, presents an invention similar to Davidson's, but using a mesh of polymeric fibers. The invention is restricted to use in repairing bone defects, and not as a bone cement in the sense described for total joint arthroplasty. Draenert, in U.S. Pat. No. 4,718,910, describes a bone cement mixture where a second phase of fibers is added. The fibers, however, are made up of the same polymeric material as the bone cement. Draenert includes a graph of the performance of the new material versus existing cements. The inventor states that the fiber is only effective because of the shape of the prepolymer powder. Therefore, the improvement is due to the use of a different cement, and not to the addition of the fibers.
Ducheyne et al. in U.S. Pat. No. 4,963,151 disclose the distribution of short, fine, reinforcing fibers homogeneously throughout surgical bone cement by adding the fibers in the form of bundles of several hundred fibers with the fibers bonded to each other with an adhesive binder that is soluble in the liquid monomer component of the bone cement.
It is generally agreed that as the quantity of reinforcing fibers increases so do the mechanical strength properties. However, as the fiber content increases it becomes increasingly difficult and eventually not practical or possible to effect homogeneous distribution of the fibers throughout the cement mass and in addition the viscosity of the mass increases and its workability by the surgeon during surgery decreases. Any practically useful surgical bone cement must be capable of being easily mixed by the surgeon in a clinical setting, i.e., during surgery, and must remain sufficiently flowable and workable to be applied to the bone surface or cavity and/or to the prosthesis or other implant device.